When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0), the individual magnetic moments of the nuclear spins in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. Usually the nuclear spins are comprised of hydrogen atoms, but other NMR active nuclei are occasionally used. A net magnetic moment Mz is produced in the direction of the polarizing field, but the randomly oriented magnetic components in the perpendicular, or transverse, plane (x-y plane) cancel one another. If, however, the substance, or tissue, is subjected to a magnetic field (excitation field B1; also referred to as the radiofrequency (RF) field) which is in the x-y plane and which is near the Larmor frequency, the net aligned moment, Mz, may be rotated, or “tipped” into the x-y plane to produce a net transverse magnetic moment Mt, which is rotating, or spinning, in the x-y plane at the Larmor frequency. The practical value of this phenomenon resides in the signal which is emitted by the excited spins after the excitation field B1 is terminated. There are a wide variety of measurement sequences in which this nuclear magnetic resonance (“NMR”) phenomenon is exploited.
When utilizing these signals to produce images, magnetic field gradients (Gx, Gy, and Gz) are employed. Typically, the region to be imaged experiences a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The emitted MR signals are detected using a receiver coil. The MRI signals are then digitized and processed to reconstruct the image using one of many well-known reconstruction techniques.
Magnetic resonance angiography (MRA) uses the NMR phenomenon to produce images of the human vasculature. There are three main categories of techniques for achieving the desired contrast for the purpose of MR angiography. The first general category is typically referred to as contrast enhanced (CE) MRA. The second general category is time-of-flight (TOF) MRA. The third general category is phase contrast (PC) MRA.
To perform CE MRA, a contrast agent, such as gadolinium, is injected into the patient prior to the magnetic resonance (MR) angiogram to enhance the diagnostic capability of the MR angiogram. While CE MRA is a highly effective means for noninvasively evaluating suspected vascular disease, the technique suffers from several additional drawbacks. First, the contrast agent that must be administered to enhance the blood vessel carries a significant financial cost. Second, contrast agents such as gadolinium have recently been shown to be causative of an often catastrophic disorder called nephrogenic systemic fibrosis (NSF). Third, CE MRA does not provide hemodynamic information, so that it is not always feasible to determine if a stenosis is hemodynamically significant. Fourth, the signal-to-noise ratio (SNR) and, therefore, spatial resolution is limited by the need to acquire data quickly during the first pass of contrast agent through a target vessel. For these reasons, there have been substantial efforts to move away from CE MRA imaging protocols in favor of non contrast-enhanced (NCE) MRA protocols.
Fortunately, TOF and PC MRA imaging techniques do not require the use of a contrast agent. The 3D TOF techniques were introduced in the 1980's and they have changed little over the last decade. The 3D TOF MRA techniques commonly used for cranial examinations and have not been replaced despite recent advances in time-resolved contrast-enhanced 3D MRA. An alternative technique known as pulsed arterial spin labeling (PASL) was first applied to image intracranial circulation years ago; however, image quality never approached that of 3D TOF and the method has had little clinical utility. Moreover, electrocardiographic (ECG) gating was required. The use of TOF MRA is generally limited to imaging of intracranial circulation, however, because of sensitivity to patient motion and flow artifacts.
Finally, phase contrast (PC) MRA is largely reserved for the measurement of flow velocities and imaging of the vasculature. Phase contrast sequences are the basis of MRA techniques utilizing the change in the phase shifts of the flowing protons in the region of interest to create an image. Spins that are moving along the direction of a magnetic field gradient receive a phase shift proportional to their velocity. Specifically, in a PC MRA pulse sequence, two data sets with different amounts of flow sensitivity are acquired. This is usually accomplished by applying gradient pairs, which sequentially dephase and then rephase spins during the sequence. The first data set is acquired using a “flow-compensated” pulse sequence or a pulse sequence without sensitivity to flow. The second data set is acquired using a pulse sequence designed to be sensitive to flow. The amount of flow sensitivity is controlled by the strength of the bipolar gradient pulse pair use in the pulse sequence because stationary tissue undergoes no effective phase change after the application of the two gradients, whereas the different spatial localization of flowing blood is subjected to the variation of the bipolar gradient. Accordingly, moving spins experience a phase shift. The raw data from the two data sets are subtracted to yield images that illustrate the phase change, which is proportional to spatial velocity.
Thus, to perform PC MRA pulse sequences, a substantial scan time is generally required and the operator must set a velocity-encoding sensitivity, which varies unpredictably depending on a variety of clinical factors. The necessity of acquiring two data sets negatively affects the temporal resolution by a factor of two. Moreover, the two acquisitions use different gradient waveforms (e.g. a flow-encoded data set uses a bipolar gradient whereas a flow-rephased data set uses three gradient lobes). Consequently, gradient-induced eddy currents are not identical for the two acquisitions, which results in spatially and time-varying background phase shifts despite the use of image subtraction. Such background phase shifts cause errors in the velocity measurement and necessitate the use of complex phase correction algorithms.
Accordingly, several attempts have been made to utilize a “referenceless” phase contrast MRI technique. For instance, Nielsen J F, Nayak K S, Referenceless phase velocity mapping using balanced SSFP. Magn. Reson. Med, 2009 May; 61(5): 1096-102, uses a balanced SSFP pulse sequence with the requirement that the echo time (TE) be equal to one-half of the repetition time (TR). The use of a bSSFP sequence is not desirable because of an intrinsic sensitivity to artifacts from off-resonance effects and increased artifacts from gradient-induced eddy currents compared with a spoiled gradient-echo acquisition. Moreover, the method requires the manual placement of regions of interest near blood vessels and additional processing with a phase correction algorithm. Even with all these steps, the background phase correction is not uniform.
Therefore, it would be desirable to have a system and method for decreasing the acquisition time of phase-based, flow-encoding, imaging techniques that does not correspondingly increase artifacts in the resulting images.